This invention relates to an improved method and apparatus for locating the origin of radiation, such as a gamma ray, in an imaging system. In particular, this invention relates to a method and apparatus for determining the origin of a gamma ray as part of a technique of imaging internal organs in a living subject.
There are many situations in medicine where it is desirable to obtain images of a patient's internal organs or body functions. The imaging technology that is used to do this includes a variety of techniques such as magnetic resonance imaging (MRI), computed tomography (CT), single photon emission computed tomography (SPECT), and positron emission tomography (PET) .
Generally speaking, in PET, a radionuclide is administered internally to a living subject. A positron from the decaying radionuclide encounters and annihilates with an electron, resulting in a pair of 511 keV annihilation photons which are emitted in exactly opposite (180.degree.) directions from the annihilation site in the subject. By arranging banks of radiation detectors--typically scintillation detectors--all around the subject, the origin of the gamma ray can be determined. Substantially simultaneous detection of photons in opposingly situated detectors defines the site of the positronelectron annihilation as lying somewhere along a line directly between the opposing detectors. Typical PET scanners or tomographs include complex computerized data systems for collecting the information obtained and using it to reconstruct an image of the target organ, using mathematical techniques similar to those employed in computerized tomography.
The radioactive isotopes used in PET include, but are not limited to, .sup.18 F, which has a half-life of approximately 110 minutes, .sup.11 C (half-life of approximately 20 minutes), .sup.13 N (half-life of approximately 10 minutes), and .sup.15 O (half-life of approximately 2 minutes). Because of the relatively short half-lives of the radioisotopes used, they are typically produced in an on-site cyclotron or other particle accelerator. Other nuclides exist which have either a long half life or a parent with long half life. These can be used without on-site cyclotrons, but they have generally less desirable chemical or physical characteristics. The practical need for an on-site cyclotron dramatically increases the cost of PET and therefore has limited the number of such systems in place.
In contrast, in SPECT, a single photon is emitted from a radionuclide at a site in the patient's body. The photon is again detected, but in contrast to PET, the origin of the photon is determined by analyzing the information obtained when the single photon strikes different portions of an array of radiation detecting elements, thereby permitting the deduction of its path. SPECT uses longer-lived isotopes than PET, including but not limited to .sup.99m Tc (half-life of approximately 6 hours) and .sup.201 T1 (half-life of approximately 74 hours). However, the resolution obtainable through SPECT imaging is lower than that presently available in PET systems.
In both prior art PET and SPECT systems, the scintillating detectors are able to detect the emitted photons (also called gammas) by means of a phenomenon whereby a photon interacts with an atom of the scintillating detector, which may be in the form of a scintillating optical fiber. This interaction results in the ejection of a so-called photoelectron or Compton electron. The ejected electron transfers energy to atomic, molecular, or crystalline structures in the fiber, and causes the emission of light quanta. The light propagates toward an end of the fiber, where it is detected by means such as photomultipliers. The ejected electron, meanwhile, will sometimes have sufficient energy to move on and interact with at least one more scintillating fiber in an array of alternating x-y planes of orthogonal fibers, again resulting in the generation of light in those fibers. By detecting the light generated in the two or more fibers, and then determining the point at which those fibers intersect, one can determine the site of the event.
It is crucial for electronically collimated SPECT imaging that the path of the incoming gamma ray be determined. This is possible by detecting two (Compton) events within the same detector or two different detectors; the line through the two points at which the fibers intersect determines the direction of the incoming gamma ray.
Presently, both PET and SPECT systems employing optical fibers have less than optimal resolution and efficiency of detection of gamma rays. Two factors that reduce the efficiency of these systems are: (1) gammas (photons) which pass through the scintillating fiber array without generating a photoelectric or Compton event, and (2) events which go undetected due to the requirement that the electron traverse at least two layers of fibers in the z direction in order for detection to be accomplished. This poses a particular problem in SPECT systems at low energy. A fiber of for example 0.25 mm diameter will stop an electron of kinetic energy 150 keV; a fiber of 0.05 mm diameter stops electrons of 50 keV. Thus Compton interactions of a few hundred keV gamma rays most frequently excite only one fiber in prior art radiation detecting systems. In such a situation, the position of the origin of the gamma ray is not detectable without using extremely fine fibers, which are both inherently inefficient and expensive.
There is a need for improved imaging methods and apparatus that would eliminate or reduce the shortcomings of the prior art, not only in PET and SPECT, but also in other imaging technologies.